Cardiac arrhythmia, such as atrial fibrillation, occurs when regions of cardiac tissue abnormally conduct electric signals to adjacent tissue, thereby disrupting the normal cardiac cycle and causing asynchronous rhythm. Important sources of undesired signals are located in various tissue regions in or near the heart, for example, the ventricles, the atria and/or and adjacent structures such as areas of the pulmonary veins. Regardless of the sources, unwanted signals are conducted abnormally through heart tissue where they can initiate and/or maintain arrhythmia.
Procedures for treating arrhythmia include surgically disrupting the origin of the signals causing the arrhythmia, as well as disrupting the conducting pathways for such signals. More recently, it has been found that by mapping the electrical properties of the heart muscle in conjunction with the heart anatomy, and selectively ablating cardiac tissue by application of energy, it is possible to cease or modify the propagation of unwanted electrical signals from one portion of the heart to another. The ablation process destroys the unwanted electrical pathways by formation of non-conducting lesions.
In this two-step procedure-mapping followed by ablation-electrical activity at points in the heart is typically sensed and measured by advancing a catheter containing one or more electrical sensors into the heart, and acquiring data at a multiplicity of points. These data are then utilized to select the target areas at which ablation is to be performed.
A typical ablation procedure involves the insertion of a catheter having a tip electrode at its distal end into a heart chamber. A reference electrode is provided, generally taped to the patient's skin. Radio frequency (RF) current is applied to the tip electrode, and flows through the surrounding media, i.e., blood and tissue, toward the reference electrode. The distribution of current depends on the amount of electrode surface in contact with the tissue, as compared to blood which has a higher conductivity than the tissue. Heating of the tissue occurs due to its electrical resistivity. If the tissue is heated sufficiently, cellular and other protein destruction ensues; this in turn forms a lesion within the heart muscle which is electrically non-conductive. During this process, heating of the electrode also occurs as a result of conduction from the heated tissue to the electrode itself. If the electrode temperature becomes sufficiently high, possibly above 50 degree C., blood clot could form on the surface of the electrode. If the temperature continues to rise, more blood clot is formed while dehydration ensues.
The tip temperature increase and the associated clot formation have two consequences: increased electrical impedance and increased probability for stroke: The former relates to clot dehydration. Because dehydrated biological material has a higher electrical resistance than heart tissue, impedance to the flow of electrical energy into the tissue also increases. Increased impedance leads to sub-optimal energy delivery to the tissue which results in inadequate lesion formation, reduced ablation efficiency and eventually to sub-optimal clinical outcome. The latter, a safety hazard, is due to possible dislodgment of the formed clot and relocation in the brain vasculature. It is therefore beneficial from a safety perspective as well as ablation efficiency to minimize the tip temperature increase and clot formation. This should be accomplished without compromising the formation of lesions of appropriate sizes.
In a typical application of RF current to the endocardium, circulating blood provides some cooling of the ablation electrode. However, there is typically a stagnant area between the electrode and tissue which is susceptible to the formation of dehydrated proteins and coagulum. As power and/or ablation time increases, the likelihood of an impedance rise also increases. As a result of this process, there has been a natural upper bound on the amount of energy which can be delivered to cardiac tissue and therefore the size of RF lesions. In clinical practice, it is desirable to reduce or eliminate impedance rises and, for certain cardiac arrhythmias, to create larger lesions. One method for accomplishing this is to monitor the temperature of the ablation electrode and to control the RF current delivered to the ablation electrode based on this temperature. If the temperature rises above a pre-selected value, the current is reduced until the temperature drops below this value. This method has reduced the number of impedance rises during cardiac ablations but has not significantly increased lesion dimensions. The results are not significantly different because this method continues to rely on the cooling effect of the blood which is dependent on the location within the heart and the orientation of the catheter to the endocardial surface.
Another method is to irrigate the ablation electrode, e.g., with physiologic saline at room temperature, to actively cool the ablation electrode instead of relying on the more passive physiological cooling provided by the blood. Additionally, due to the irrigation-mediated dilution of blood around the tip, the probability for clot creation is further reduced. Thus, irrigation tip cooling and blood dilution allow for safer increase of applied RF power. This results in lesions which tend to be larger usually measuring about 10 to 12 mm in depth.
The clinical effectiveness of irrigating the ablation electrode is dependent upon the distribution of flow within and around the surface of the tip electrode structure as well as the rate of irrigation flow through the tip. Effectiveness is achieved by reducing the overall electrode temperature and eliminating hot spots in the ablation electrode which can initiate coagulum formation. More channels and higher flows are more effective in reducing overall temperature and temperature variations, i.e., hot spots. Irrigation is utilized during the entire time the catheter resides inside the patient's body. Higher flow rate is used during ablation while lower-maintenance-flow rate is required in order to prevent back flow of blood into the coolant passages during non-ablation time. The coolant flow rate must be balanced against the amount of fluid that can be safely injected into the patient. Thus, reducing coolant flow by utilizing it as efficiently as possible is a desirable design objective.
One method for designing an ablation electrode which efficiently utilizes coolant flow is the use of a porous material structure. Such designs have the advantage of distributing the coolant evenly across the entire electrode structure. This balanced cooling results in a) eradication of possible surface or interior hot spots, and b) uniform dilution of blood at the vicinity of the electrode, thus further minimizing the chance for clot formation. Such designs are described in U.S. Pat. Nos. 6,405,078 and 6,466,818 to Moaddeb et al., the entire disclosures of which are incorporated herein by reference. Moaddeb describes the use of sintered metal particles to create a porous tip electrode. In addition, Moaddeb uses a non-conductive insert implanted into the porous tip electrode for mounting a thermocouple, lead wire and/or irrigation tube within the porous tip electrode. However, during irrigation the sintered metal particles can disintegrate and break away from the electrode structure. This-undesirable-particle dislodgement may be further facilitated during ablation. Additionally, (and in the context of our MRI compatibility claims-see below) the metallic material proposed for such porous tip is not optimal for MRI imaging. Furthermore, the proposed tip does not allow for high density mapping-a highly desired feature for accurate arrhythmia diagnosis. Consequently, a desire arises for a porous electrode having increased structural integrity, being compatible with the MRI environment, and allowing for high mapping density.
A porous tip electrode catheter is also described in U.S. Pat. No. 8,262,653 to Plaza. The porous tip electrode comprises a porous material through which fluid can pass. The porous tip electrode is covered with a thin coating of conductive metal having openings (pores) through which fluids can pass. However, the porosity of such thin conductive coating is not easily controlled leading to inconsistent pore size and distribution. Therefore, distribution of irrigation fluid around the tip electrode may not be even or uniform. Furthermore, in this design, RF power delivery is achieved via direct connection (e.g. by soldering or other similar technique) of the RF power line to the tip's outer conductive coat. Thus, the presence of the generally non-uniform porous coating is necessary in order to establish electrical contact of the tip to the heart tissue.
Safe and efficacious ablation depends not only on optimal irrigation arrangement for the tip but also on accurate mapping of the electrophysiological behavior of the heart, which would allow for accurate diagnosis and appropriate tissue targeting. The greater the accuracy of the mapping the more accurate the diagnosis and thus the effectiveness of treatment. Improved (high resolution) cardiac mapping requires the use of a multitude of electrodes in close proximity to sense electrical activity within a small area, for example, a square centimeter or less.
Metallization of ceramics is a well-established technique and is widely used in a multitude of electronics and engineering disciplines, including fabrication of RF electronic circuits. Metallization involves the application of metal on ceramic substrates, including the formation of conductive regions, such as metallized conductor patterns or uniform metal layers on surfaces of ceramic substrates. Common ceramic substrates include aluminum oxide, beryllium oxide, ferrite, barium titanate, as well as quartz or borosilicate. Generally, ceramic metallization processes fall into three categories: thin-film, thick-film, and co-firing techniques. In the thin film approach, a thin layer of metal is deposited by vacuum processes such as sputtering, evaporation, chemical vapor deposition, and laser ablation. Electroless and electrolytic plating are also frequently grouped in the thin film category. To enhance adhesion, a preliminary adhesion-promoting layer, such as chromium or titanium, is often deposited. Thick film methods involve printing metal pastes, typically metal powders mixed with glass fits and organic binders onto ceramic substrates. The printed substrates are fired to form conductive paths on the ceramic. In the co-firing approach, unfired “green” ceramic surfaces are coated with patterned metal paste lines. The printed green ceramic is fired both to sinter the material and form the conductive metal patterns. Metallization processes are described, for example, in U.S. Pat. No. 4,574,094 to DeLuca, et al.; U.S. Pat. No. 5,096,749 to Harada, et al.; U.S. Pat. No. 5,690,805 to Thorn, et al.; and U.S. Pat. No. 5,725,938 Jin, et al. Metallization depending on the type of metallization process and the substrate may include gold, platinum, or other biocompatible metals suitable for intracardial signal acquisitions.
While ablation has revolutionized the treatment of cardiac arrhythmias, ablation can be improved where physicians can assess lesions in real time. The use of magnetic resonance imaging (MRI) during an ablation procedure could enable physicians to assess lesions in real time. However, the ablation catheter and other associated accessory equipment can interfere with the imaging process, causing local distortions in the MRI scans. Use of appropriate MRI compatible materials is necessary to minimize these image distortions. Safety experts have cleared some metals for use during MRIs, including titanium, cobalt-chromium, copper, selected stainless steel alloys. Non-ferromagnetic metals are also MRI compatible. Such materials include copper, brass, silver, gold, aluminum, lead, magnesium, platinum and tungsten. Ceramic materials as well as other thermoplastic polymers are non-metallic and as such are highly desirable as MRI compatible materials. They not only present minimal image distortion but being electrical insulators they present no heating effects due to absence of internally induced electrical currents. Ceramic materials of porous construction are proposed in the current invention as materials for the construction of the catheter's tip.
In view of the foregoing, it is desirable to provide a catheter with a dome tip electrode made of a porous substrate for more uniform irrigation, where the dome tip electrode incorporates surface electrodes made via a metallization, printing or other process for any desirable surface electrode pattern that provides multiple electrodes in close proximity for high density mapping. It is also desirable to provide a catheter where the substrate and the surface electrodes are MRI compatible so that the physician can conduct lesion assessment in real time during an ablation procedure.